Method and device for using impedance measurements based on electrical energy of the heart

ABSTRACT

A method and a device are disclosed for evaluating the cardio-circulatory and pulmonary condition of a patient, including determining the patient&#39;s thoracic impedance based on information solely derived from the electrical energy generated by the patient&#39;s own heart.

CROSS-REFERENCE TO RELATED APPLICATION

This application is a continuation-in-part of Ser. No. 10/622,184 filedJul. 16, 2003, now abandoned which is a continuation-in-part of Ser. No.10/155,771 filed May 25, 2002 now U.S. Pat. No. 6,829,503 that claimspriority of German patent 10148440-2 filed Oct. 1, 2001 of the sameapplicant, each of which applications is incorporated by reference inits entirety herein. Applicant claims priority of the aforesaidapplications with respect to common subject matter.

BACKGROUND OF THE INVENTION

The present invention relates generally to obtaining a measure of apatient's thoracic impedance, and more particularly to doing so usinginformation derived from an EKG signal with a device that relies onelectrical energy provided by the patient's own heart.

Specific resistance of biological materials and impedance measurementshave played a major role in modern medicine. The electrical conductivityand capacity of disperse systems have been described as early as 1931(Fricke et al., The Electric Conductivity of and Capacity of DispersedSystems; Physics 1931; 1:106-115). Later, especially in the 1950s and1960s, significant interest was directed towards the resistance ofbiological materials (e.g., Geddes L. et al., The Specific Resistance ofBiological Material: A Compendium of Data for the Biomedical Engineerand Physiologist, Medical and Biological Engineering 1967, 5:271-293).The application of impedance and resistance measurements forcardio-circulatory function by measuring the blood and body temperaturehas been studied extensively by Geddes et al., Medical and BiologicalEngineering 1967, 11:336-339). Also, internal and external whole bodyimpedance measurements have been used for noninvasive monitoring anddetermination of cardiac output (Carter et al., Chest 2004,125:1431-1440). In addition, the feasibility of using intracardiacimpedance measurements has been evaluated by E. Alt et al. for capturedetection in connection with cardiac pacing (Pace 1992, 15:1873-1879).

Background patents that describe the use of impedance in conjunctionwith implantable devices are referenced in U.S. Pat. No. 5,003,976 toAlt, which describes the cardiac and pulmonary physiological analysisvia intracardiac measurements with a single sensor. The '976 patentdiscloses that a single functional parameter, namely intracardiacimpedance, varies both with the intrathoracic pressure fluctuationsfollowing respirations and with cardiac contraction. This value isrepresentative of both pulmonary activity and cardiac activity. Thefinding indicates that this information derived from intracardiacimpedance can be used not only to monitor the patient's cardiac andpulmonary activity, condition and cardio-circulatory status, but also,to control the variability of the rate of an implantable cardiacpacemaker.

U.S. Pat. No. 4,884,576 to Alt et al. discloses a self-adjusting rateresponsive cardiac pacemaker and method based on the intracardiac signalderived from impedance measurements using an electrode implanted intothe heart. And U.S. Pat. No. 4,919,136, also to Alt, describes aventilation controlled pacemaker which uses the ventilation signalderived from those impedance measurements with an electrode in the heartto adjust the pacing rate.

Recently, considerable interest has been focused on the monitoring ofcongestive heart failure by means of impedance. U.S. Pat. No. 6,473,640to Erlebacher describes a system that detects changes in resistance to aflow of current in the systemic venous system, and detects changes inimpedance to a flow of current through lungs. The specific signalprocessing enables a determination of congestion in the venous or in thepulmonary system by application of differential signal processing ofimpedance. Other methods, such as are described by Combs in U.S. Pat.No. 5,957,861 and Riff in U.S. Pat. No. 5,876,353, respectively pertainto impedance monitoring for discerning edema through evaluation ofrespiratory rate, and use of implantable medical devices for measuringtime varying physiological conditions, especially edema, and forresponding thereto.

U.S. Pat. No. 6,104,949 to Pitts-Crick relates to a device and a methodused for the diagnosis and treatment of congestive heart failure. Godie,in U.S. Pat. No. 6,351,667, describes an apparatus for detectingpericardial effusion, in which a wire probe anchored to the right heartventricle and two other wire probes are used to measure the impedancebetween the different probes in order to assess the degree ofpericardial effusion.

U.S. Pat. No. 4,899,758 to Finklestein et al describes a method andapparatus for monitoring and diagnosing hypertension and congestiveheart failure. U.S. Pat. No. 6,336,903 to Brody relates to an automaticsystem and method for diagnosing and monitoring congestive heart failureand the outcomes thereof. U.S. patent publication 2002-0115939 toMoligan et al describes an implantable medical device for monitoringcongestive heart failure in which incremental changes in parameter dataover time provide insight to the patient's heart failure state.

The measurement of heart failure becomes of greater clinical interestand importance as more than 5 million patients in the U.S. are affected.With deterioration of myocardial function, patients often requirerepeated hospitalization. Current methods of monitoring congestive heartfailure cannot reliably predict an early occurrence of this congestiveheart failure; but an understanding of its occurrence may provide anearly indicator of this adverse event for the patient.

A considerable number of new treatment forms have been introduced intoclinical practice. It had been shown that congestive heart failure canbe treated, not only by drugs, especially Beta blockers, but also bybiventricular pacing. This method makes use of the exact timing of astimulus, not only to the right ventricle or to the septum, but also tothe left side of the heart by means of an electrode which is implantedinto the coronary venous circulation. By these means, the left ventriclecan be stimulated at a time that provides an optimal synchronization ofthe heart and improves the mechanical effectiveness of the systole bysynchronizing the depolarization of the right heart, the septum and theleft heart. This avoids the ineffective late contraction of the leftventricle at a time when the septum depolarization has already occurred,and the squeezing of the blood by synchronous action of the septum andleft ventricle is no longer present. In addition, the reduction inmitral valve regurgitation by this type of resynchronization has beenshown.

Studies published at the 2005 meeting of the American College ofCardiology in Orlando, Fla., USA (CARE-HF study) illustrate that notonly the quality of life of those patients with New York HeartAssociation, Heart Failure Class 3 and 4 can be improved, but also thelife expectancy. This recent data show very impressively that over a3-year period such biventricular stimulation and the mortality can bereduced by half in a highly significant manner. All these new devicesimprove the survival and quality of life of patients and have abeneficial effect on re-hospitalization. Nevertheless, the occurrence ofheart failure is still a major problem for these patients, and it isbeneficial to detect such a heart failure as early as practicable.

A second parameter which plays a major role in patients with implantabledevices such as pacemakers and defibrillators is the correct adjustmentof heart rate. Rate adaptive pacemakers in the past have provided anopen type of correlation between a signal parameter to adjust the heartrate and the affected heart rate. However, even multiple sensorparameters that have been used for adjustment of the pacing rate havenot brought the real need of a patient to clinical practice, mainly aclosed-loop monitoring of heart rate.

In the healthy person, the heart rate is regulated by a verysophisticated closed loop and negative feedback. Heart rate onlyincreases to a level with exercise which is physiologically beneficial.This means that if a patient exercises only mildly, his heart rateincreases proportional to the increase in oxygen uptake for this personwhich is a fraction of his maximum exercise capacity, maximum oxygenuptake and aerobic and anaerobic capacity. Thus, if someone iswell-trained, an external load of 50 watts might represent only 25% ofhis/her maximum exercise capacity if the patient is capable ofexercising up to a level of 200 watts. With this external load of 50watt the heart rate will increase by only the fraction that isrepresented by the patient's resting heart rate and maximum exerciseheart rate. In other words, such a well-trained person will increasehis/her heart rate only by 30-35 beats per minute (bpm). A less capablepatient who has a maximum exercise capacity of 100 watts, will increasehis/her heart rate with the same external load to a higher degree. Inthat case, the slope of increase in heart rate depends not only on afixed relation of a sensor parameter, such as ventilation or physicalactivity or any other physiologic parameter having a suitablecorrelation with heart rate, but also on his/her underlyingcardio-pulmonary exercise capacity and condition.

It is therefore a principal aim of the present invention to provide anovel method to detect a parameter that can control not only the heartrate in a physiologic appropriate manner, and provide a closed-loopfeedback control for the optimal heart rate of an exercising patient,but also to monitor the individual status of the patient under a widerange of physiologic conditions including normal resting status,congestive heart failure and exercise states.

Many attempts have been made in the past to use impedance measurementsto derive appropriate signals; however, the past approaches haveinvolved use of external power sources to power the device(s) that wouldmonitor and detect impedance. This external energy can be applied eitheroutside the thorax from a supply external to the body, or by animplantable device that uses energy from an electrical battery housedwithin the device itself.

SUMMARY OF THE INVENTION

It is the aim of the present invention to provide means and methods tomonitor impedance of a patient by using the patient's own heart as thepower source.

It is a further aim of the invention to provide a method of detectingthe thoracic impedance of a person without need for a battery or otherexternal power source or need for the respective circuitry to providethe current or voltage for the impedance measurement that has heretoforelimited the availability of energy to power implantable devices, andaccordingly required periodic and even relatively frequent replacementof the implanted device.

It has been long known that the EKG, which can be derived form thesurface of the patient, represents a voltage generated by the heart.This voltage is derived from the skin of a patient by means ofelectrodes which are attached. The resulting voltage in an EKG can bedetected from different leads. There are bipolar electrodes which derivea voltage, between for example Lead I the right arm and the left arm, inthe way that the resulting voltage change between these two electrodesrepresents the main vector of the heart in projection to those leads.Therefore, the amplitude is a measurement of the voltage generated bythe heart and the vector. The input impedance of an external EKG machineis standard in a range between 1-10 megohms. This means that the inputimpedance and resistance of such an amplifier is very high and thereforeno current is shunting through the machine and the voltage alwaysrepresents the maximum voltage generated from the energy source, mainlythe heart. Differences in voltages with the current EKG measurementresult from different vectors that project two different leads on thesurface of a patient.

The same holds true for voltages detected with implantable devices fromleads which are situated within the heart or within the thorax or evenimplantable devices which have EKG electrodes outside the thorax, suchas EKG loop recorders or devices which are suitable for monitoring theEKG and congestive heart failure from electrodes that are situatedoutside the thoracic cage such as described by Alt et al in theaforementioned related '771 patent application.

The underlying principle of the invention may be summarized, forexemplary purposes, from experiments conducted by the applicants. Themeasurements that resulted from placement of standard EKG I, II, and IIIleads on the patient were recorded in the presence and absence of anexternal load. The amplitude of the EKG signal that corresponds to themeasured voltage is a function of the impedance of the EKG amplifier.

The underlying theory of the invention is that the heart acts as abattery. A battery fails when its internal resistance has increased to alevel at which the battery can no longer supply a useful amount of powerto an external load. That same principle applies to the measurement ofelectrical energy generated by the heart. That is, if several loads areapplied to the measurement device, which may be an implantable cardiacpacemaker, a defibrillator, a device for monitoring the occurrence ofheart failure, or a diagnostic device for monitoring the physicalcondition of a patient, the same phenomena can be used to calculate theinternal impedance at the site of measurement. Preferably, thecalculation or determination is of the thoracic (and preferably,intrathoracic) impedance or of local impedance and/or its relativechanges with time for a given patient.

According to an important aspect of the invention, a method ofevaluating the cardio-circulatory condition of a patient would includedetermining the patient's thoracic impedance based on information solelyderived from the electrical energy generated by the patient's own heart.The intrathoracic impedance information may be used, for example, tooptimize the rate response of a rate adaptive pacemaker, or to optimizemonitoring of the patient's congestive heart failure. Another aspect ofthe invention, then, may be characterized as a method of adjusting theheart rate of a patient by means of an implanted rate adaptivepacemaker, where automatic adjustment of the pacing rate of thepacemaker is achieved in response to a determination of the patient'sintrinsic impedance derived solely from the electrical energy generatedby the patient's heart. The patient's intrinsic impedance informationmay be used instantaneously to influence the rate adaptation on acontinuous or ongoing basis within minutes. Alternatively, the rate andcardio-pulmonary response derived from the intrinsic impedance may beapplied to determine the rate adaptation on a longer term basis, such ason a daily or monthly basis.

It is therefore another important aim of the invention to provide asystem which uses an impedance derived parameter, such as ventilation,as a closed loop parameter to optimize the rate response of animplantable rate adaptive cardiac pacemaker or defibrillator.

The invention may be further stated as allowing the cardio-pulmonarystatus of a patient to be monitored with an implantable monitoringdevice, by calculating the patient's thoracic impedance based oninformation derived from the electrical energy generated by thepatient's own heart. For example, the device may be implantedsubcutaneously to monitor the patient's EKG, and to detect changes inthe thoracic impedance based on differential signal processing of theEKG. Then, information concerning the impedance changes may be appliedwithin the device to determine the cardio-pulmonary status of thepatient. In circumstances where the patient is suffering from heartfailure, the device is adapted to monitor the patient's heart failure byperforming the calculation of impedance and processing thereof solelyusing the electrical energy generated by the patient's own heart.

Information about the cardiac function of the patient may be obtainedusing a method according to the invention, in which electrical signalinformation from an EKG derived from depolarization and repolarizationof the patient's heart, representing systole and diastole, accordinglyfor different phases of the heart represented by the EKG, iscontinuously processed using electrical energy generated by thepatient's own heart. On the other hand, the impedance of the patient'sheart may be analyzed with systole from a point close to the T-Wave ofthe EKG signal, and information on the diastolic status of the heart maybe derived from an impedance signal at the R-Wave of the EKG signal.Then, a comparison between systole and diastole may be used to ascertainindirectly cardiac stroke volume and the cardio-pulmonary status of thepatient.

Similarly, the function of a body-implantable defibrillator may beenhanced according to the invention by determining the impedance betweensensing electrodes of the defibrillator, and determining changes in thatimpedance, based solely on energy generated by the patient's heart.

It will be seen from the ensuing detailed description that a device forevaluating the cardio-circulatory condition of a patient is implementedwith means for determining the patient's thoracic impedance andimpedance changes based on information solely derived from theelectrical energy generated by the patient's own heart. And the desiredinformation may be obtained by extremely simple means so that thechanges or additions required to achieve these benefits with evencurrently available devices can be minimal, such as including surfacemounted electrodes for monitoring the patient's surface EKG insubcutaneously implanted devices.

BRIEF DESCRIPTION OF THE DRAWINGS

The above and other aims, objectives, aspects, features and advantagesof the invention will be better understood from a consideration of thefollowing detailed description of the best mode contemplated forpracticing the invention, taken with reference to certain preferredimplementations and methods, and the accompanying drawings in which:

FIG. 1A-1D are schematic diagrams of an electrical circuit or system inwhich a patient's heart is represented by an internal resistance, usefulfor explaining the basic principle of the present invention;

FIG. 2 applies the concepts discussed with reference to FIG. 1, to theprinciple of different EKG vectors, FIG. 2A illustrating EKG vectorsmeasured for a patient, and FIG. 2B illustrating the EKG tracings fordifferent placements of EKG electrodes;

FIG. 3 illustrates various situations in which the internal resistanceof a patient's heart is represented in parts A, B, C and D not as onesingle value, but by plural individual impedances;

FIG. 4 is a graph of EKG data and impedance factor versus time thatillustrates results from measurements taken with an intracardiacelectrode;

FIGS. 5 and 6 are graphs of EKG data and impedance factor vs. timecorresponding to FIG. 4, except for changes (respective increases) inthe respiratory rate;

FIG. 7 is a graph that summarizes and compares the measurements ofimpedance derived ventilatory rate obtained using two different methods;

FIG. 8 is a graph that depicts a preferred application of the signalprocessing means, with with relative signal amplitudes over time;

FIG. 9 is a graph that represents the quotient of raw voltage and shuntvoltage with different depth and modes of breathing;

FIG. 10 is a graph that depicts adjustment of pacing rate, in units ofpacing rate versus extent of a patient's physical activity detected by asensor; for rate responsive pacing;

FIG. 11 is a flow chart that shows the principle of a closed loop rateprofile optimization; and

FIG. 12 is a flow diagram of the principles and signal processing of thepresent invention.

DETAILED DESCRIPTION OF THE PRESENTLY CONTEMPLATED BEST MODE OFPRACTICING THE INVENTION

The method and device of the invention will be better understood byreference to a presently preferred embodiment constituting a stand alonediagnostic or therapeutic device to enhance the specificity of abody-implantable device such as an artificial cardiac pacemaker, adefibrillator, a cardiac resynchronization device, or a monitoringdevice for evaluating the cardiopulmonary functional status of thepatient.

Throughout the views of the drawings, identical reference numbersindicate identical structures. Views of the device or the method eitheralone or implanted are not intended to represent the actual or relativesizes or rather give principle understanding of the methods and devices.

FIGS. 1A-1D are simplified schematic diagrams of an electrical circuitor system in which the heart of a patient 1 is represented by aninternal resistance (R_(internal)) 2, at least in parts 1A and 1B. Inprinciple, R_(internal) is the sum of many individual resistances orimpedances consisting of myocardial tissue, fibrous connective tissue inthe heart, pericardium, blood within the heart, fluids within theintracellular spaces, and the surrounding environment of connectivetissue atria, pulmonary structures, venous structures, and lung tissue.The multiplicity of cells within the heart depolarize, and thisdepolarization creates an electrical force which threads through theheart with a certain vector. This initial electrical source representedin the circuit schematic of the Figures as a voltage source E (3) has amagnitude that can be measured between electrodes at points A (4) and B(5). The latter two points may constitute the sites of electrodes of anexternal measuring instrument or EKG amplifier.

Virtually the same voltage as the original source voltage E 3 will bepresent and remain so, if the impedance between measuring points 4 and 5has a magnitude of several megohms. This is because under that conditionlittle or no current is flowing between those points. In principle, thismaximum voltage is present, for example, at the input amplifier of animplantable pacemaker, since they have an impedance of several hundredkilohms or megohms; and the same is true of other implantable diagnosticdevices, such as devices that measure the patient's EKG. For example,external EKG strip chart recorders or EKG monitors have an impedance of1 megohm or more, thereby allowing them to detect the maximum voltagepresent between points of the body at which their electrodes areattached or located.

In FIG. 1B the schematic shows a heart 1, in circuit with an internalresistance 2, a source voltage 3 that is detected between electrodepoints or sites 4 and 5, corresponding to what has been described abovefor FIG. 1A. However, in this case an additional external resistance ofmagnitude R_(Load) 6 is connected in circuit between electrode points 4and 5. If the magnitude of the resistance of load 6 is considerablylower than the input impedance between 4 and 5 represented by, say, thevirtual open circuit impedance of an amplifier, discussed for thecircuit of FIG. 1A, considerable current will be shunted through thislower resistance. As in the case of a failing battery, then, the initialfull voltage E (3) will not be detected since, by Ohm's Law (U=IR), adrop in voltage will be observed with decreased R.

This example is carried forward in FIGS. 1C and 1D. In each of thoseFigures, the heart 1 generates a certain EKG signal 7 between measuringpoints 4 and 5, which has a certain amplitude 8. In FIG. 1D, theconnection of an input amplifier 8 with a relatively low resistanceR_(Load) 6 across points A and B results in an EKG signal 7 having anamplitude considerably lower than that in the case of the much higher,virtually open circuit resistance across A and B in FIG. 1C.

In principle this observation can be compared to a battery. When abattery fails it is typically because its internal resistance hasincreased to a level that no longer supports the supply of a usefulamount of electrical energy to an external load. If one measures thevoltage V of a failing battery which is disconnected, it is usuallyfound that the battery has a nearly normal voltage because aconventional voltmeter used to perform the measurement has an inputresistance much higher than the internal resistance of the battery. If,however, the failing battery is connected to a low external resistancesuch as load 6 in FIG. 1B, the terminal voltage U of the battery dropsprecipitously. This can be interpreted as the battery dropping most ifits source voltage across its own internal resistance, so little or novoltage is available for external services. For example, an idealbattery with 0 internal resistance or infinite internal conductance anda voltage E of 12 volts, when supplying power to an external load havinga resistance of 1 ohm, will produce a current of I=12 amps and a powerof E×I=144 watts. If the battery has an internal resistance of 2 ohm, oran internal conductance as low as 0.5 siemens, then with this load, theterminal voltage U of the battery will drop to 4 volts. The outputcurrent of this failing battery will drop to I=U/R_(Load) which is 4amps, and the output power is 16 watts.

The same principle holds for conventional electrocardiography along themain electrical excitation vector. Since an EKG measurement is detectedwith high input impedance, this conventional measurement gives noinsight into the electrical power of the source, in this case the heartof the patient. Furthermore, the absolute voltage of EKG signals is nota valuable indicator of various pathological situations. Indeed, despitegreat diversity of cardiac diseases it is common clinical experiencethat individual variability and the amplitude of the EKG wave asdetected from state of the art EKG amplifiers is not indicative of anykind of disease. The voltage of a conventional EKG is reduced only invery few clinical situations when large electrical shunts are present,such as a pericardial effusion which constitutes a large conductoraround the heart that shunts the electrical energy with low intrinsicresistance.

The terminal voltage U in FIG. 1B (which represents the R wave amplitude8, represented in FIG. 1C by QRS-complex 7) should drop byE−(R_(internal)×I)), where I=U/R_(Load). The internal resistanceR_(internal) 2 can be calculated by the equationR_(internal)=(V−U)/(U/R_(Load)), where V is the voltage betweenelectrode points 4 and 5 which is disconnected from the load in FIG. 1Aand U is the terminal voltage between 4 and 5, across which theelectrical load R_(Load) 6 is connected. Based on R_(internal), theinternal electrical conductance S_(internal) of the heart can also becalculated by the equation:S _(internal)=1/R _(internal), in siemens.

These equations are applied in practice in FIG. 2, according to theprinciple of different EKG vectors. FIG. 2A represents a patient 9 inwhich EKG lead I 10 is represented by electrodes to the patient's rightand left arms, with voltage measurements shown in EKG tracing 11 (FIG.2B). Electrode detection lead/electrodes II 12 produces the tracing 13,and the voltage detected between the electrodes of lead III 14 is shownin the tracing 15 for lead III. Tracings 11, 13, and 15, then, representthe EKG signal obtained for the respective vector projections of theleads 10, 12 and 14. If an external load is now connected in parallelwith the input impedance of the EKG amplifier, a voltage drop will beobserved for the same patient as shown by tracings 16 (for lead I), 17(for lead II), and 18 (for lead III), because part of the energydelivered from the patient's heart is shunted through the additionalexternal load.

FIG. 3 illustrates various situations in which the internal resistanceR_(internal) of the heart is represented not as one single value, but byplural individual impedances. In FIG. 3A, the internal resistanceR_(heart) of the heart 19 is represented by the structural resistance ofthe heart made up of cells, connective tissue, and primarily solids, anda variable component of resistance R_(inspiration) (or R_(insp)) 20 isprimarily represented by the filling of the heart with blood. Sinceblood has a specific impedance of roughly 50 ohms per centimeter (cm),while the specific impedance of heart 19 is 400 ohms per cm, thereexists a great influence on the total impedance of the heart, becausethese two components are in parallel. The internal voltage source 3 inFIG. 3A detected between electrode points 4 and 5 represents a voltageV₁ (22) that equals primarily E 3 if the input impedance between 4 and 5is sufficiently high that all of the voltage E drop occurs between 4 and5.

If, however, an additional external load R_(load) 6 of roughly 1 kilohmor less is applied as shown in FIG. 3B, then the voltage between 4 and 5drops to a voltage U1 (23) which, as earlier described, is lower thanthe voltage E 22. FIG. 3C illustrates the situation in which a variationnow occurs in internal resistance 21, representing the resistance withexpiration (R_(expiration) or R_(exp)). The total voltage to be detectedV₂ 24 is now primarily composed of the parallel resistances of R_(heart)19 and R_(exp) 21. A variation between R_(insp) 20 and R_(exp) 21 willnot affect the resulting voltages, V₂ 24 or V₁ 22, since the inputimpedance between electrodes 4 and 5 is sufficiently high to avoidfurther voltage shunting and voltage drop. However, as shown in FIG. 3D,the external load 6 will affect voltage U₂ 25 with a variation ininternal impedance 21 during expiration, compared to impedance component20 during inspiration. Thus, if an external load 6 of sufficient loadresistance, such as 1 kilohm, is applied to a primarily high inputimpedance amplifier, variations in internal total resistance build upfrom R_(heart) 19 and R_(respiration) 20 or 21 have a much greatereffect on voltage U₁ 23 with inspiration and U₂ 25 with expiration. Inprinciple, for this condition it can be said that R_(inspiration) is notidentical with R_(expiration) and therefore, U₁ 23 is different from U₂25. It follows that U₁ equals the delta of U₂, and this represents moreor less the impedance factor of respiration, the term “impedance factor”meaning the quotient of impedances 19 and 20 in FIG. 3B compared to theimpedances 19 and 21 in FIG. 3D.

FIG. 4 is a graph of EKG data and impedance factor versus time thatillustrates results from measurements taken with an intracardiacelectrode. A bipolar conventional pacemaker electrode was implanted inthe heart and measurements were taken between the electrode tip inconnection with the myocardium and a ring located roughly 1 cm behindthe electrode tip. These sites can be considered as electrode points 4and 5 in the Figures described thus far, and a linear high qualityamplifier was connected between these two sites. The signal processingwas performed in such a way that one signal represented in FIG. 4 as rawvoltage 26 represented by the higher bars in the graph was compared to ashunt voltage 27 represented by the smaller bars. To detect the shuntvoltage from the same electrode site 4 and 5 by a special program, theinput impedance was shunted by a resistance of one kilohm. In FIG. 4 thetime axis (abscissa) 28 shows increments of time in seconds and thevoltage axis (left ordinate) 29 shows increments of the detected voltageof the two signals raw voltage 26 and shunt voltage 27. The curve 30represents the quotient between voltage 26 and voltage 27 (i.e., theirimpedance factor, measured along lines parallel to the right ordinate)or in other words, the quotient of the impedances that change withrespiration.

As is clearly seen in the graph, the ratio of the peak signal betweenraw voltage 26 and shunt voltage 27 represented by curve 30 correlateswith the respiration, which was set to 5.5 cycles per minute. The timeinterval for one respiratory cycle is 11 seconds in this example, whichactually represents a ventilation rate of 5.5 cycles per minute.

Various aspects of the continuous EKG signal can be used to derivemeasurements of impedance factor in FIG. 4, to discern or determine thecardio-pulmonary status of the patient using, in this example, thebipolar conventional pacemaker electrode implanted in the heart formonitoring purposes. Either a continuous line can be averaged if asufficiently high digitization rate is applied, or, to simplifymeasurements and procedures, and also to facilitate data handling andpower consumption in an implantable device, only certain aspects of theEKG signal need be taken. For example, the latter aspects may be thoserepresented previously herein in EKG signal 7 with amplitude 8, so it isfeasible to use only the peak 7A of the R Wave or to take other aspectssuch as only or additionally the T-Wave peak 7B, of the EKG signal 7illustrated in FIG. 2B. In the example shown in FIG. 4, the peak of theR-Wave was applied. From the latter Figure, it is clear thatconsiderable variation occurs in the quotient represented by curve 30(the impedance factor) between inspiration and expiration, whichcorresponds to the ventilatory cycle rate and its amplitude.

FIGS. 5 and 6 are graphs illustrating the same data setting and the sameparameters as in FIG. 4; however, the respiratory rate was changed inFIG. 5 to 10 cycles, and in FIG. 6 to 20 cycles per minute. This changein frequency is clearly shown in the latter two Figures, beingrepresented by peak ratio 30.

FIG. 7 is a graph that summarizes and compares the measurements of trueventilatory rate V along the abscissa and of the impedance factorderived ventilatory rate Z along the ordinate. As shown, there is nearlya 1:1 correlation between the two methods which means that ventilationcan be truly detected, as far as the frequency of breaths is concerned,from the impedance derived signal in which the heart serves as the powersource for the impedance calculation.

FIG. 8 depicts a preferred application of the signal processing means,with the abscissa showing units of time and the ordinate showing thesignal amplitudes of the raw voltage 26 and the shunt voltage 27relative to one another. The former is indicated by the higher amplitudeand the latter, derived from the shunted input impedance which wasreduced to 0.5 kilohm, is indicated by a lower amplitude. The signalprocessing of both signals 26 and 27 included passage through a low passfilter for smoothing.

The application of low pass and high pass filtering in the signalprocessing can be performed by either a conventional analog technique oralso by conventional digital technique of first, second, third or fourthorder. The selection of the cut off frequency depends on preference asto which signal components should be detected. If it is preferred todetect respiration and respiration rate by filtering with a cut offfrequency below twice the maximum frequency that is expected. For mostpatients, a respiratory rate of no more than 50 breaths per minute canbe expected, which means that a low pass filter of about 1.5 Hz or lesswill allow detecting the respiratory rate. In FIG. 8 the difference insignal amplitude between signal 26, which was detected from inputimpedance of more than 1 megohm, and signal 27 detected from an inputimpedance of 0.5 kilohm, becomes evident from FIG. 8. The respiratorycycle rate is also evident.

FIG. 9 is a graph that represents the quotient of raw voltage and shuntvoltage with different depths or amplitudes and modes of breathing. Thisgraph shows that it is not only feasible to detect respiratory rate,which is effected by change in the filling of the heart with blood andby a volume change of the amount of blood surrounding an electrodeimplanted in the heart with a conductor such as blood, but it is alsofeasible to detect a relative change in amplitude following differenttidal volumes. Graph 31 represents the signal derived from the impedancequotient of high input impedance of more than 1 megohm and 1 kilohm withexternal artificial ventilation of an individual with a tidal volume of300 ml per breath. Line 32 shows the same for a tidal volume of 850 ml.Line 33 depicts the impedance quotient with spontaneous breathing at aconsiderably lower rate. Now the signal shows not only the breathing orventilation, but also the cardiac component 33A indirectly reflectingstroke volume with systole and diastole. In order to obtain this, forexample, one need only observe the depolarization that occurs with theelectrical signal represented by the peak of the R wave 7A in the EKG orrelative to the repolarization that occurs with the peak of the T wave7B, recognizing that the mechanical contraction occurs slightly afterthe peak of the R wave. At this point, intracardiac impedance allowsdetermining the extent of filling of the heart and to derive therefromindications of heart failure.

FIG. 10 is a graph that depicts, in units of pacing rate versus extentof physical activity detected by a sensor, the principle currently usedfor rate responsive (or rate adaptive) pacing in cardiac pacemakers andin defibrillators, to adjust the heart rate or the pacing rate of apatient based on the input received from a sensor (such as anaccelerometer in heretofore available devices). In the Figure, sensorunits 35, which may, for example, be activity counters, amplitude of anactivity signal, or any other physiological parameter which has acorrelation to heart rate in healthy persons such as body temperature,ventilation, or mechanical forces acting on the body. These units aretranslated in a slope function into a rate response with more sensorunits yielding a high pacemaker rate 34. Based on empirical assumptionsa certain slope is selected for a given patient. Slope A (36) representsa correlation of sensor units and pacemaker rate for a physically fitpatient, while a slope B (37) is selected for a less physical patient.On average there is a threshold on which the heart rate is increaseduntil maximum exertion 38 and maximum heart rate 34 are reached.

The limitation with all these open loop systems of the prior art is thatthe slope that sets an individual rate with a given exercise is selectedon an empirical basis; however, it is not confirmed that this heart rateis the optimum heart rate for the given situation of a patient. Even inthe same patient, the most beneficial rate with a given exercise mightchange from day to day. Some patients have significant coronary heartdisease which limits the flow of blood through the coronary arteries andtherefore induces an ischemia Therefore, it is sometimes beneficial tolimit the maximum heart rate and the slope to different values (e.g.,39) compared to a state that might have been present weeks or even onlydays ago when the myocardial perfusion was different. The lack of afeedback parameter is one of the limitations of currently widespread useof rate adaptive pacemakers.

Considering the limitation of currently open loop rate regulationimplantable pace makers, an impedance derived parameter according to theinvention can be used not only to adjust the pacing rate directly so asto create a correlation between signal and pacing rate on a linear ornonlinear basis as has been suggested in the past by ventilation controlrate adaptive pacemakers, but it also can be used to control theindividually optimum pacing and heart rate on a closed loop basis, andalso on a long term trend basis. In the past it has been considered thathemodynamic parameters would be suitable for a closed loop system.However, the limitation is to derive hemodynamic parameters directly.The complexity and change in those parameters and the technicaldifficulty to measure an actual derivative of cardiac output or strokevolume has prevented an introduction of those practices into clinicalpractice. In stark contrast, the present invention provides a novelminimally energy demanding system that allows controlling the effect ofpacing rate and the adequacy of pacing rate by monitoring theventilation and/or cardiac parameters derived from an impedance signalwhich is obtained in a manner according to the invention.

FIG. 11 is a flow chart that shows the principle of a closed loop rateprofile optimization. In this optimization, the impedance derived ratecontrolled parameter is processed and stored in circuitry 40 which wouldprocess the signals obtained and store them in a conventional memorymeans of an implantable device such as an implantable pacemaker ordefibrillator. This can be done either as a long term average overseveral days, or over a shorter term average over a daily or even anhourly or minute-by-minute basis. The difference between long term andshort term averages as described in previous patents by the applicantsreferenced earlier herein may be applied to this process. The actualrate profile is available from the second circuitry 41. This actual rateprofile is compared to the long term impedance derived rate controlparameter in a comparator and logic circuit 42, and the output is usedto adjust the heart rate setting at 43. This not only controls theslope, for the correlation of pacing rate to a given signal intensity orwork load as shown in the graph 44, but may also be used to set a newtarget rate and base rate. In this way, an optimization of the rateprofile and slope can be achieved by means of minimal additionalhardware complexities, since the EKG electrodes, the EKG amplifier andthe information is already available to any demand pacemaker anddefibrillator.

Alternatively, the signal can be also processed on a short-term, so thatthe effect of the heart rate adjustment can be evaluated against theventilation and cardiac response within tens of seconds to achieve anindividual optimization by circuitry 42. The pacing rate that gives thelowest ventilation response is the optimum for a given work load. It isknown that a person who cannot increase his heart rate will breathe moreheavily than one who has the more adequate heart rate. Experiments doneby the applicants have shown in the past that if the pacing rate is keptconstant at 70 bpm and there is no further rate increase with exercise,the exercise capacity of a patient is limited and he/she will breatheheavily despite the limited exercise capacity. Therefore, the responseof ventilation can be taken as an indirect indicator of the metabolicload and of the efficiency of the cardio-circulatory system. In thisway, the invention fulfills an aim to provide a system that usesventilation as a closed loop parameter to optimize the rate responsegiven brought forward by a different sense of parameter. In addition,ventilation can also be used as a primary parameter to adjust the heartrate with or without secondary optimization provided through anothersense parameter.

FIG. 12 is a flow diagram of the principles and signal processing of thepresent invention. An EKG signal sensor 45 consists of electronic meansthat makes connection with the site where the EKG signal is sensedeither merely on an intracardiac basis, from a bipolar electrode, from aunipolar electrode (one electrode in the heart and one on the case), orfrom separate electrodes on an implantable case such as surface-mountedelectrodes on a pacemaker, a defibrillator, or monitoring device that ispreferably implanted subcutaneously, or implanted in a differentlocation in the body as described in the cross-referenced relatedapplications. At least bipolar EKG signals are sensed, and by aswitching relay process 46 a load 6 is either added or not added to thesignal circuitry. Thereby, two different EKG derived information signalsare available for EKG amplifier and processor 47, one signal with andone signal without the load. Comparator and analyzer 48 derives thedesired information from the quotient of the two impedances thatcorrespond to voltages and impedances derived from a high and a lowinput impedance. This information is then provided on either a filteredor processed basis to storage medium 49 in which several long term,short term, and derivatives of the signals can be stored. Comparator 50finally derives the information that is required to the respectiveimplantable device either for a rate profile adjustment and optimization51, or if it is only for monitoring purposes, for congestive heartfailure 52 as described as one of exemplary embodiments in theaforementioned related '184 patent application. This telemonitoringsignal can be telemetered to an outside device or through telemetry 53or can also provide patient alert 54 in case of detecting a deviationfrom the desired pattern for the individual patient. Electronic meansapplied are state of the art, with switching relay process, EKG signalsensor, amplifier processor, analyzer, storage and comparator providedeither in a single chip or the like, or in conventional electroniccomponents applying a combination of digital and analog techniques or insolely digital techniques including filtering and effecter 55 thatconverts the available information into the desired action within thedevice.

In summary, the current invention provides a facilitated monitoringmeans to optimize both therapeutic and diagnostic capacity of animplantable device by means of impedance derived information, and toacquire this information using the patient's own heart to provide thenecessary electrical energy.

Although a presently contemplated best mode of practicing the inventionhas been disclosed by reference to certain preferred methods, it will beapparent to those skilled in the art from a consideration of theforegoing description that variations and modifications may be madewithout departing from the spirit and scope of the invention.Accordingly, it is intended that the invention shall be limited only bythe appended claims and the rules and principles of applicable law.

1. A method of evaluating a condition of a patient having a heart and athoracic impedance, comprising the steps of: detecting an electricalsignal from said heart of said patient using signal circuitry viaelectrode means arranged and adapted for sensing said electrical signal,while using electrical energy generated by said heart of said patientand no other electrical power for said signal circuitry; and determiningsaid thoracic impedance based on said electrical signal.
 2. The methodas in claim 1 further comprising the step of optimizing a function of arate adaptive pacemaker based, at least in part, of said thoracicimpedance.
 3. The method as in claim 1 further comprising the step ofoptimizing monitoring of a congestive heart failure of said patientbased, at least in part, of said thoracic impedance.
 4. The method as inclaim 1 wherein said electrical signal is at least partiallyrepresentative of an EKG of said patient.
 5. The method as in claim 4wherein said EKG is sensed on an intracardiac basis.
 6. The method as inclaim 5 wherein said detecting step is accomplished, at least in part,by a bipolar electrode.
 7. The method as in claim 5 wherein saiddetecting step is accomplished, at least in part, by a unipolarelectrode.
 8. The method as in claim 7 wherein said detecting step isaccomplished, at least in part, with a device adapted to be implanted insaid patient, wherein said detecting step utilizes one electrodepositioned in said heart of said patient and electrically connected tocircuitry within an implanted device and another electrode constitutingan electrically conductive portion of a case of said device.
 9. Themethod as in claim 4 wherein said detecting step utilizes a deviceadapted to be implanted in said patient and utilizes surface mountedelectrodes associated with said device.
 10. The method as in claim 4wherein said detecting step senses from said electrical signal two EKGderived information signals reflecting relatively high and low inputimpedance loads on said electrode means.
 11. The method as in claim 10wherein said determining step processes said two EKG derived informationsignals to obtain information indicative of short-term and long-termtrends of said thoracic impedance of said patient.
 12. The method as inclaim 1 wherein said thoracic impedance comprises impedance frommyocardial tissue, fibrous connective tissue in the heart, pericardium,blood within the heart, fluids within the intracellular spaces, and thesurrounding environment of connective tissue atria, pulmonarystructures, venous structures, and lung tissue.
 13. A method ofadjusting a heart rate of a patient having a heart and thoracicimpedance by means of an implantable rate adaptive pacemaker having apacing rate, comprising the steps of: determining said thoracicimpedance from said heart of said patient using signal circuitry whileusing electrical energy generated by said heart of said patient and noother electrical power for said signal circuitry; adjusting said pacingrate of said implantable rate adaptive pacemaker in response to saidthoracic impedance.
 14. The method as in claim 13 further comprising thestep optimizing said pacing rate through closed loop optimization. 15.The method as in claim 13 further comprising the step of adapting aheart rate of a rate adaptive pacemaker based, at least in part, on saidthoracic impedance.
 16. The method as in claim 13 further comprising thestep of instantaneously influencing adaptation of said pacing rate on anongoing basis within minutes based, at least in part, on said intrinsicimpedance.
 17. The method as in claim 13 further comprising the step ofdetermining said adaptation of said pacing rate on a long term dailyand/or monthly basis based, at least in part, on long term rate andcardio-circulatory response from said intrinsic impedance information ofsaid patient.
 18. A method of monitoring a cardio-circulatory status ofa patient having a heart and a thoracic impedance with an implantabledevice, comprising the steps of: detecting an electrical signal fromsaid heart of said patient using signal circuitry while using electricalenergy generated by said heart of said patient and no other electricalpower for said signal circuitry; and calculating said impedance based onsaid electrical signal.
 19. The method as in claim 18 wherein saiddetecting step is accomplished subcutaneously.
 20. The method as inclaim 18 wherein said electrical signal is at least partiallyrepresentative of an EKG of said patient and wherein said calculatingstep is at least partially accomplished through differential signalprocessing of said electrical signal.
 21. The method as in claim 20further comprising the step of applying information concerning changesin said impedance to determine said cardio-circulatory status of saidpatient.
 22. The method as in claim 18, wherein the patient is sufferingfrom heart failure, further comprising the step of monitoring said heartfailure by performing said calculating step in calculating saidimpedance.
 23. A method of enhancing function of a body-implantabledefibrillator having sensing electrodes adapted to function inconjunction with a heart of a patient, comprising the steps of:detecting an electrical signal generated by said heart of said patientusing signal circuitry while using electrical energy generated by saidheart of said patient and no other electrical power for said signalcircuitry; and determining an impedance between said sensing electrodesof said defibrillator based, at least in part, on said electricalsignal; and determining changes in said impedance between said sensingelectrodes of said defibrillator based, at least in part, on saidelectrical signal.
 24. A device for monitoring a cardio-circulatorystatus of a patient having a heart and a thoracic impedance with animplantable device comprising: signal circuitry means for detecting anelectrical signal generated by said heart of said patient while usingelectrical energy generated by said heart of said patient and no otherelectrical power for said signal circuitry means; and means forcalculating said thoracic impedance based on said electrical signal. 25.A method of obtaining information about a cardiac function of a patienthaving a heart, comprising the steps of: detecting an electrical signalgenerated by said heart of said patient related to an EKG using signalcircuitry while using electrical energy generated by said heart of saidpatient and no other electrical power for said signal circuitry; andcontinuously processing said electrical signal during depolarization andrepolarization of said heart of said patient, said electrical signalrepresenting systole and diastole as different phases of said heart. 26.The method as in claim 25 further comprising the steps of: analyzing animpedance of said heart and changes of said impedance with systole froma point close to a T-Wave of said electrical signal; and derivinginformation on a diastolic status of said heart derived from saidelectrical signal close to an R-Wave of said electrical signal.
 27. Themethod as in claim 26 further comprising the step of comparing saidimpedance at systole with said impedance at diastole to assess acardio-circulatory status of said patient.
 28. An implantable medicaldevice for monitoring a cardio-pulmonary status of a patient having aheart and a thoracic impedance, comprising: signal circuitry means fordetecting an electrical signal generated by said heart of said patientrelated to an EKG while using electrical energy generated by said heartof said patient and no other electrical power for said signal circuitrymeans; and means for determining said thoracic impedance based on saidelectrical signal related to said EKG.
 29. The implantable medicaldevice of claim 28, wherein said implantable medical device is adaptedfor subcutaneous implantation.
 30. An implantable device as in claim 28for monitoring a cardio-pulmonary status of a patient having a heart,comprising means for determining a thoracic impedance of said patientfrom an EKG signal information acquired using electrical energygenerated by said heart of said patient; and surface mounted electrodes,operatively coupled to said means for determining, for sensing said EKGsignal information; wherein said device is adapted for subcutaneousimplantation.
 31. The implantable medical device as in claim 28 furthercomprising signal processing means, operatively coupled to said meansfor determining, for deriving information concerning changes in saidthoracic impedance to determine the cardio-pulmonary status of thepatient.
 32. The device of claim 28, wherein said implantable medicaldevice is adapted for cardiac pacing, and further comprising means fornegative feedback closed loop control of said cardiac pacing foroptimization of a pacing rate to match a fitness level of said patientengaged in physical activity.